1. Field of Invention
The field of this invention relates to surgical guidewires, and more particularly to surgical guidewires for use with Interventional Magnetic Resonance Imaging.
2. Discussion of Related Art
The contents of all references, including articles, published patent applications and patents referred to anywhere in this specification are hereby incorporated by reference.
Interventional Magnetic Resonance Imaging (“iMRI”) has increased in interest during the last decade due to the advent pulse sequence improvements, Magnetic Resonance (“MR”) compatible instruments, the development of rapid imaging techniques and automatic instrument tracking techniques.
For MR guidance of vascular interventions to be safe, the interventionalist must be able to visualize the tip location and distal shaft of the MRI compatible guidewire relative to the vascular system and surrounding anatomy. A number of instrument visualization approaches under MRI have been developed including both passive and active techniques. Passive visualization techniques rely on the creation of susceptibility artifacts to enhance the device (e.g., catheter) appearance by using contrast agents or ferromagnetic materials. Active visualization relies on supplemental hardware embedded into a catheter body such as a Radio Frequency (“RF”) antenna to receive the RF signal during MRI (Susil R C, Yeung C J, Atalar E, “Intravascular extended sensitivity (IVES) MR imaging antennas.” Magnetic Resonance in Medicine, 2003; 50(2): 383-390). However, none of these techniques provides satisfactory results in terms of both instrument tip and shaft visualization at the same time. Visualization of the shaft only is not enough to advance a guidewire through tortuous vessels due to the risk of puncturing vessel walls and visualization of a single point is not sufficient for steering an active guidewire in complex vessel territory (Atalar E, Kraitchman D L, Carkhuff B, Lesho J, Ocali O, Solaiyappan M, Guttman M A, Charles H K, Jr. Catheter-tracking FOV MR Fluoroscopy. Magnetic Resonance in Medicine 1998; 40(6):865-872).
Interventional MR Imaging
Magnetic Resonance Imaging (MRI) is one of the most important clinical imaging modalities. A significant advantage of using MRI in clinical procedures is that imaging via MR is conducted only using a strong homogenous magnetic field and radio frequency energy pulses, without the use of harmful ionizing radiation such as with the use of X-ray angiography. Also, MRI utilizes Nuclear Magnetic Resonance principles with gradient coil elements to provide spatial encoding, resulting in the ability to perform 3-D human body imaging with high soft tissue contrast (Lauterbur P C. NMR Imaging in Biomedicine. Cell Biophysics 1986; 9 (1-2): 211-214; Lai C M, Lauterbur P C. True Three-Dimensional Image Reconstruction by Nuclear Magnetic Resonance Zeugmatography. Physics in Medicine and Biology 1981; 26(5):851-856; Kramer D M, Schneider J s, Rudin A M, Lauterbur P C. True Three-Dimensional Nuclear Magnetic Resonance Zeugmatographic Images of a Human Brain. Neuroradiology 1981; 21(5):239-244). MRI allows one to obtain information about various physical parameters such as flow, motion, magnetic susceptibility and temperature (Axel L. Blood Flow Effects in Magnetic Resonance Imaging. Magnetic resonance Annual 1986; 237-244; Henkelman R M, Stainsz G J, Graham S J. Magnetization Transfer in MRI: A review. NMR Biomedicine 2001; 14(2):57-64; Dickenson R J, Hall A S, Hind A J, Young I R. Measurement of Changes in Tissue Temperature using MR Imaging. Journal of Computer Assisted Tomography 1986; 10(3):468-472). Because of this, MRI has a wide variety of both diagnostic and therapeutic imaging applications both in the clinical and research environment. When MRI was initially introduced in the clinical environment, it was used for only diagnostic imaging purposes with almost no consideration for use in therapeutic procedures (Webb W R, Gamsu G, Stark D D, Moon K L, Jr., Moore E H. Evaluation of Magnetic Resonance Sequences in Imaging Mediastinal Tumors. American Journal of Roentgenology 1984; 143(4):723-727; Belli P, Romani M, Magistrelli A, Masetti R, Pastore G, Costantini M. Diagnostic Imaging of Breast Implants: Role of MRI. Rays 2002; 27(4):259-277). Reasons for this can be attributed to the lack of sequences designed for interventional MRI such as sequences for real-time device tracking, sequences that provide image contrast that correlate directly to therapy performance, high-speed sequences that allow real-time imaging with sufficient contrast and resolution and the lack of dedicated and optimized hardware for interventional applications.
In recent years, efforts have been made to develop MRI as an interventional tool for image guided interventions by addressing the above mentioned challenges (Miles K. Diagnostic and Therapeutic Impact of MRI. Clinical Radiology 2002; 57(3):231-232; Jolesz F A, Blumenfeld S M. Interventional Use of Magnetic Resonance Imaging. Magnetic Resonance Quarterly 1994; 10(2):85-96; Jolesz F A. Interventional and Intraoperative MRI: A General Overview of the Field. Journal of Magnetic Resonance Imaging 1998; 8(1):3-7). Also, development of 1.5 T magnets with short bores that allow access to the groin area for catheter-based procedures, liquid crystal image displays that can be exposed to high magnetic fields, improvements in the hardware of the magnetic field gradient systems for additional gains in image acquisition speed, the development of catheter based MRI antennas for localized intravascular signal reception provide wide range of interventional MR Imaging applications.
Mri Compatible and Visible Devices for Interventional Procedures
MR guided interventions can only be performed with devices free of ferromagnetic components, otherwise as one would encounter severe magnetic forces (induced displacement force and torque) on the device by the static magnetic field of the MR scanner and they would also cause image distortion (Shunk K A, lima J A, Heldman A W Transesophageal magnetic resonant imaging. Magn Reson. Med 1999; 41:722-726). However, MR compatible and safe devices are not enough to perform vascular interventions with MRI. The reliable visualization of these devices in relation to the surrounding tissue morphology is also required. In contrast to ultrasound, X-ray fluoroscopy, or computed tomography (CT), visualization of interventional instruments in MR has proven to be difficult. A number of approaches have been developed for depicting vascular instruments in an MR environment. They can be broadly grouped into two categories: passive and active visualization.
Passive Visualization
In passive visualization techniques, achieving adequate catheter contrast is based on enhancing the inherent signal void of an instrument as it displaces spins during insertion. Differences in magnetic susceptibility can be used to create large local losses in signal due to intra-voxel dephasing (Rubin D L, Ratner A V, Young S W. Magnetic susceptibility effects and their application in the development of new ferromagnetic catheters for magnetic resonance imaging. Invest radiol. 1990; 25:1325-1332). The tip or body of passive catheters is composed of either ferromagnetic or paramagnetic sleeves that produce susceptibility artifacts. Incorporating multiple rings of paramagnetic dysprosium oxide (Dy2O3) along the instrument tip allows the catheter to be consistently visualized independently of orientation (see FIG. 1) (Bakker C J, Hoogeveen R M, Hudak W F, van Vaals J J, Viergever M A, Mali W P. MR-guided endovascular interventions:susceptibility-based catheter and near real time imaging technique).
Susceptibility markers must have a high magnetic moment to induce an adequate artifact at a variety of scan techniques and tracking speeds. In other words, they must have sufficient contrast to noise ratio (CNR) with respect to the background in order to distinguish the device in thick slab images. The CNR can be calculated as
                    CNR        =                  -                                    S              m                                      σ              s                                                          (        1        )            
With Sm the signal loss induced by the marker, which is negative in the subtraction image, and σs is the noise in the background of the subtraction image. Because an MR scanner provides a strong magnetic field with applied gradients, the passive marker material is magnetized homogenously. So, the magnetic moment (m) can be directly related to B0 (Teitz J R, Miltford F J and Christy R W 1993 Foundations of Electromagnetic Theory 4th edn (Reading, M A: Addison-Wesley) p245). This relation can be written asm(B0)=ΔM(B0)V  (2)
with ΔM the magnetization difference between the marker and tissue and V the volume of the marker. For diamagnetic and paramagnetic materials, m is proportional to B0. For other types of magnetism like ferro- and ferrimagnetism, a nonlinear relation exists. For a marker of subvoxel volume, the field distortion can be approximated by a dipole field (Bos C, Viergever M A and Bakker c J G 2003 On the artifact of a subvoxel susceptibility deviation in spoiled gradient-echo imaging Magn. Reson. Med. 50 400-4)
                              Δ          ⁢                                          ⁢                                    B              z                        ⁡                          (                              x                ,                y                ,                z                            )                                      =                                                            μ                0                            ⁢                              m                ⁡                                  (                                      B                    0                                    )                                                                    4              ⁢              π                                ⁢                                                    2                ⁢                                  z                  2                                            -                              y                2                            -                              x                2                                                                    (                                                      x                    2                                    +                                      y                    2                                    +                                      z                    2                                                  )                                            5                2                                                                        (        3        )            
with μ0 the permeability in free space and x, y and z the spatial coordinates with ΔBz and B0 in the z-direction. The magnetic moment of the marker needs to be large with regard to the size of the marker and CNR of the artifacts. This excludes diamagnetic and weakly paramagnetic materials, since their magnetization is too low to create a large magnetic moment with only a small amount of material (Schenck J F 1996 the role of magnetic susceptibility in magnetic resonance imaging: MRI magnetic compatibility of the first and second kinds Med. Phys. 23 815-50). High tracking speeds can be achieved when short echo times are used, and this requires an even higher magnetic moment for heavy T2*-weighting around the marker. Then, strongly paramagnetic markers become bulky in order to realize an adequate magnetic moment. Biocompatibility and safety can be achieved by embedding into another medium, e.g. the coating of the device. Again, easily magnetizable materials are preferred, since a lower concentration of marker material is needed. The optimal applicability regarding field strength may be realized with a field-strength-independent magnetic moment of the marker. Such a nonlinear magnetic property is partly exhibited by ferro- and ferromagnetic materials. The advantage of using a passive marker is that circuit components and transmission lines are not required to visualize the catheter. This property of passive visualization techniques is really important because it also eliminates electrical safety issues. However, this technique also has several disadvantages. First of all, it provides low spatial resolution. Secondly, it slows down the speed of the procedure compared to active tracking methods. And finally, a susceptibility artifact varies based on device orientation and magnetic field strength.
Active Visualization
Active visualization relies on the incorporation of a miniature solenoid coil into the device itself (Dumoulin C L, Souza Sp, Darrow R D. Real-time position monitoring of invasive devices using magnetic resonance. Magn reson. Med. 1993; 29:411-415; Ladd M E, Zimmerman G G, Mcklnnon G C, von Schulthess G K et al. Visualization of vascular guidewires using MR tracking. J Magn Reson Imaging 1998; 8:251-253; Leung D A, debatin J F, Wildermuth S, McKinnon G C et al. Intravascular MR Tracking catheter:preliminary experimental evaluation. Am J roentgenol 1995; 164:1265-1270). The coil is connected to the scanner via a thin coaxial cable passing through the catheter and provides a robust signal, identifying the instrument location with high contrast. The tip of an active catheter can be visualized with high contrast by the incorporated coil on the tip.
Loop Antenna: Solenoid Coil
A solenoid coil is basic form of loop antenna element in which the wire is coiled in a helical pattern to create a cylindrical shape as shown in FIG. 2. Solenoid micro coils can be connected to the MR systems through the use of coaxial transmission lines, which may serve both detuning and signal transduction purposes. Loop antenna signal sensitivity for small-loop receivers falls off very rapidly (1/r3, where r is the radial distance from the loop) (Balanis C A. Antenna theory. New York: John Wiley & Sons; 1997. 941 p.). To improve longitudinal coverage, long-loop intravascular antennas were subsequently investigated (Atalar E, Bottomley P A, Ocali O, Correia L C, Kelemen M D, Lima J A, Zerhouni E A. High resolution intravascular MRI and MRS by using a catheter receiver coil. Magn Reson Med 1996; 36:596-605). For these long, narrow loop receivers (in which the loop length is much greater than its width), sensitivity falls off as 1/r2. Despite improved longitudinal coverage, these receivers produce limited SNR when conductor separation is small as can be seen in FIG. 3.
Loop Antenna: Opposed Solenoid Coil
The opposed solenoid loop antenna is based on groups of helical loops separated by a gap region, with current driven in opposite directions in the helical loops on either side of the gap, as shown in the schematic of FIG. 4. The gap provides the small area of homogenous longitudinal magnetic field that makes it a good candidate for especially using as an imaging coil within and beyond the vessel wall. However, it has a small area of homogenous longitudinal coverage compared with the dipole antenna.
Dipole Antenna
A dipole antenna for iMRI applications can be a simple coaxial transmission line with an extended inner conductor. Dipole antenna sensitivity falls off as 1/r where r is the radial distance from the antenna center (Susil R C, Yeung C J, Atalar E, “Intravascular extended sensitivity (IVES) MR imaging antennas.” Magnetic Resonance in Medicine, 2003; 50(2): 383-390).
Dipole antenna sensitivity can be improved by increasing the insulation layer (insulation broadens the SNR distribution) and helical winding over the extended core inductor (winding allows for improved SNR near the tip of the antenna). A dipole antenna appearance in a water filled phantom under MR Imaging can be seen in FIG. 5.